U.S. patent number 9,283,299 [Application Number 14/298,170] was granted by the patent office on 2016-03-15 for injectable hydrogels.
This patent grant is currently assigned to WILLIAM MARSH RICE UNIVERSITY. The grantee listed for this patent is WILLIAM MARSH RICE UNIVERSITY. Invention is credited to Kristel W. M. Boere, Adam K. Ekenseair, F. Kurtis Kasper, Antonios G. Mikos, Tyler J. Touchet, Tiffany N. Vo.
United States Patent |
9,283,299 |
Mikos , et al. |
March 15, 2016 |
Injectable hydrogels
Abstract
The present disclosure generally relates to injectable
compositions. More particularly, the present disclosure relates to
injectable, thermogelling hydrogels and associated methods. In one
embodiment, the present disclosure provides for a composition
comprising a poly(N-isopropylacrylamide)-based macromer and a
polyamidoamine-based macromer.
Inventors: |
Mikos; Antonios G. (Houston,
TX), Kasper; F. Kurtis (Houston, TX), Ekenseair; Adam
K. (Boston, MA), Vo; Tiffany N. (Houston, TX), Boere;
Kristel W. M. (Utrecht, NL), Touchet; Tyler J.
(Cypress, TX) |
Applicant: |
Name |
City |
State |
Country |
Type |
WILLIAM MARSH RICE UNIVERSITY |
Houston |
TX |
US |
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Assignee: |
WILLIAM MARSH RICE UNIVERSITY
(Houston, TX)
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Family
ID: |
48575076 |
Appl.
No.: |
14/298,170 |
Filed: |
June 6, 2014 |
Prior Publication Data
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Document
Identifier |
Publication Date |
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US 20150079020 A1 |
Mar 19, 2015 |
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Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
Issue Date |
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PCT/US2012/068810 |
Dec 10, 2012 |
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61569091 |
Dec 9, 2011 |
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61642210 |
May 3, 2012 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61K
47/18 (20130101); A61L 27/52 (20130101); A61L
27/54 (20130101); A61L 27/18 (20130101); A61K
9/06 (20130101); C08L 33/26 (20130101); A61K
35/28 (20130101); A61K 38/18 (20130101); A61K
9/0024 (20130101); A61L 2300/414 (20130101); A61L
2400/06 (20130101); A61L 2430/34 (20130101); A61L
2300/252 (20130101); A61L 2300/64 (20130101) |
Current International
Class: |
A61L
27/18 (20060101); A61L 27/54 (20060101); C08L
33/26 (20060101); A61K 9/06 (20060101); A61K
47/18 (20060101); A61K 9/00 (20060101); A61K
35/28 (20150101); A61K 38/18 (20060101); A61L
27/52 (20060101) |
References Cited
[Referenced By]
U.S. Patent Documents
Primary Examiner: Azpuru; Carlos
Assistant Examiner: Hagopian; Casey
Attorney, Agent or Firm: Reed Smith LLP Riddle; Robert R.
Gibson; Matthew S.
Government Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
This invention was made with government support under Grant No. R01
DE17441 awarded by National Institute of Health. The government has
certain rights in the invention.
Parent Case Text
CROSS-REFERENCE TO RELATED APPLICATIONS
This application is a continuation-in-part of PCT/US2012/068810
filed Dec. 10, 2012 which claims priority to U.S. Provisional
Patent Application Ser. No. 61/569,091 filed Dec. 9, 2011 and U.S.
Provisional Patent Application Ser. No. 61/642,210 filed May 3,
2012, both of which are incorporated herein by reference.
Claims
What is claimed is:
1. A composition comprising a poly(N-isopropylacrylamide)-based
macromer and a diamine-based macromer, wherein the
poly(N-isopropylacrylamide)-based macromer has a polymer backbone
comprising glycidyl methacrylate (GMA) and N-isopropylacrylamide
(NiPAAm), and wherein the poly(N-isopropylacrylamide)-based
macromer is cross-linked by the diamine-based macromer via the GMA
of the poly(N-isopropylacrylamide)-based macromer and an amine
functional group of the diamine-based macromer.
2. The composition of claim 1 wherein the
poly(N-isopropylacrylamide)-based macromer is selected from the
group consisting of p(NiPAAm-co-GMA),
p(NiPAAm-co-GMA-co-DBA-co-AA), and p(NiPAAm-co-GMA-co-AAm).
3. The composition of claim 1 wherein the diamine-based macromer is
a polyamidoamine macromer.
4. The composition of claim 3 wherein the polyamidoamine macromer
comprises bisacrylamide.
5. The composition of claim 1 wherein the diamine-based macromer
increases hydrophilicity of the composition.
6. The composition of claim 1 wherein the composition is
liquid.
7. The composition of claim 1 further comprising mesenchymal stem
cells.
8. The composition of claim 1 further comprising growth
factors.
9. The composition of claim 1 further comprising mesenchymal stem
cells and growth factors.
10. A method comprising: combining a
poly(N-isopropylacrylamide)-based macromer with a diamine-based
macromer to form a composition, wherein the
poly(N-isopropylacrylamide)-based macromer has a polymer backbone
comprising glycidyl methacrylate (GMA) and N-isopropylacrylamide
(NiPAAm); and injecting the composition into a defect in a mammal
immediately following combining the
poly(N-isopropylacrylamide)-based macromer with a diamine-based
macromer.
11. The method of claim 10 wherein the defect is a craniofacial
defect.
12. The method of claim 10 wherein the poly(N-isopropylacrylamide)
macromer is selected from the group consisting of p(NiPAAm-co-GMA),
p(NiPAAm-co-GMA-co-DBA-co-AA), and p(NiPAAm-co-GMA-co-AAm).
13. The method of claim 10 wherein the wherein the diamine-based
macromer is a polyamidoamine macromer.
14. The method of claim 13 wherein the polyamidoamine macromer
comprises bisacrylamide.
15. The method of claim 10 wherein the composition does not impede
tissue formation within the defect.
16. The method of claim 10 wherein the diamine-based macromer
increases hydrophilicity of the composition and mitigates
syneresis.
Description
BACKGROUND
The majority of effort in tissue regeneration research has been
focused on implantable scaffolds. However, there are many
applications that may be better served with injectable, in situ
forming materials capable of co-delivering cells and growth factors
to optimize tissue regeneration.
In addition, current tissue engineering strategies remain limited
in providing functional and aesthetic reconstruction for complex
craniofacial trauma.
It is therefore desirable to develop an injectable scaffold system
capable of delivering in situ forming materials capable of
co-delivering cells and growth factors to optimize tissue
regeneration. It would also be desirable to develop a tissue
engineering strategy that is especially suited for providing
functional and aesthetic reconstruction for complex craniofacial
trauma.
SUMMARY
The present disclosure generally relates to injectable
compositions. More particularly, the present disclosure relates to
injectable, thermogelling hydrogels and associated methods.
In one embodiment, the present disclosure provides for a
composition comprising a poly(N-isopropylacrylamide)-based macromer
and a polyamidoamine-based macromer.
In another embodiment, the present disclosure provides for a method
comprising providing a composition comprising a
polyamidoamine-based macromer and a
poly(N-isopropylacrylamide)-based macromer; and allowing the
composition to crosslink.
In another embodiment, the present disclosure provides for a method
comprising providing a composition comprising a
polyamidoamine-based macromer and a
poly(N-isopropylacrylamide)-based macromer; injecting the
composition into a defect in a mammal; and allowing the composition
to solidify, wherein the composition solidifies at the site of the
defect.
In another embodiment, the present disclosure provides for an
injectable scaffold comprising a poly(N-isopropylacrylamide)-based
macromer and a polyamidoamine-based macromer.
In another embodiment, the present disclosure provides for a
hydrogel composition comprising a polyamidoamine-based
macromer.
In another embodiment, the present disclosure provides for a
hydrogel composition comprising a poly(N-isopropylacrylamide)-based
macromer.
The features and advantages of the present disclosure will be
apparent to those skilled in the art. While numerous changes may be
made by those skilled in the art, such changes are within the
spirit of the invention.
DRAWINGS
The patent or application file contains at least one drawing
executed in color. Copies of this patent or patent application
publication with color drawing(s) will be provided by the Office
upon request and payment of the necessary fee. Some specific
example embodiments of the disclosure may be understood by
referring, in part, to the following description and the
accompanying drawings.
FIG. 1 is drawing that illustrates a system for injectable tissues
engineering scaffolds.
FIG. 2 is a graph that illustrates the LCST concept.
FIGS. 3A, 3B, and 3C illustrate LCST modulation. FIG. 3A
illustrates the hydrolysis of DBA. FIG. 3B illustrates initial
(.tangle-solidup.) and final (.box-solid.) LCST of polymer with
different DBA mol content in complete hydrolysis conditions (pH=10
at 70.degree. C.). FIG. 3C illustrates LCST change as a function of
time of polymers with different DBA mol content during degradation
(pH=7.4 at 70.degree. C.).
FIG. 4 is a drawing that illustrates one example of the basic
structure of a thermogelling macromer of the present
disclosure.
FIG. 5 is illustrating synthesis of polyamidoamine, wherein n may
be greater than or equal to 1.
FIG. 6 is a drawing that illustrates how hydrogels may form in TGM
containing GMA with a PAMAM crosslinker.
FIG. 7 is a drawing that illustrates the results of a factorial
study design.
FIGS. 8A, 8B, and 8C illustrate gels using thermoresponsive,
chemically crosslinked macromers. TGM without GMA undergoes
syneresis without PAMAM (left vial) and remains liquid with PAMAM
(right vial) after 1 day (FIG. 8A). TGM with GMA and PAMAM
thermogelled after 1 minute (FIG. 8B) and fully formed after 30
minutes (FIG. 8C).
FIG. 9 illustrates degradation studies. Hydrolysis-dependent peak
LCST of TGM without GMA over 28 days
FIG. 10 is an image of a P-BA gel.
FIG. 11 illustrates swelling ratios of P(NiPAAm-co-GMA). Swelling
ratios of P(NiPAAm-co-GMA) at 10 and 20 wt % in PBS with P-2.6 kDa
at varying crosslinker:polymer functionality ratios.
FIG. 12 illustrates degradation rates of formulations with varying
PAMAM molecular weight. Accelerated degradation of P(NiPAAmco-GMA)
at 10 and 20 wt % in PBS pH 10 with P-1800 and at 10 wt % with
P-2600.
FIG. 13 illustrates a polyamidoamine.
FIGS. 14A and 14B illustrate the synthesis of a thermogelling
macromer (TGM).
FIG. 15 illustrates the synthesis of a polyamidoamine (PAMAM).
FIG. 16 illustrates a TGM/PAMAM gelation factorial study.
FIG. 17 illustrates gelation and equilibrium swelling.
FIG. 18 illustrates factorial study results.
FIG. 19 illustrates PAMAM degradation.
FIG. 20 illustrates PAMAM degradation.
FIG. 21 illustrates a PAMA-BA system.
FIG. 22 illustrates the results of degree of modification.
FIG. 23 illustrates PAMAM-BA.
FIG. 24 illustrates .sup.1H NMR spectra of p(NiPAAm-co-GMA) with
proton peak locations identified.
FIG. 25 illustrates DSC thermograms for pNiPAAm and
p(NiPAAm-co-GMA) at 10 wt % polymer in pH 7.4 PBS solutions.
FIG. 26 shows .sup.1H NMR spectra of PAMAM with proton peak
locations identified and equations utilized to calculate an
experimental average molecular weight based on peak intensities
displayed.
FIG. 27 shows raw microTOF mass spectroscopy data for
purified/dried PAMAM crosslinker showing molecular weights and
charge states for some of the largest peaks, with a close-up view
of a single peak family (insert).
FIG. 28 shows desired PAMAM diamine product M.sub.n, and weight
fraction of the total reaction products, as determined by microTOF
mass spectroscopy, and reaction media pH (insert) over the course
of the synthesis reaction.
FIG. 29 shows a schematic representation of the hydrogel epoxy
crosslinking reaction.
FIG. 30 shows hydrogel formation and syneresis with 10 wt % TGM and
varying PAMAM content (0-10 wt %) with volume swelling ratio
relative to 0% PAMAM, Q.sub.r, shown.
FIG. 31 shows a DSC thermogram monitoring heat flow over time
during the epoxy crosslinking reaction of a 10 wt % TGM and 7 wt %
PAMAM solution in pH 7.4 PBS.
FIG. 32 shows oscillatory rheology traces showing shear storage,
G', and loss, G'', moduli for an injectable hydrogel solution with
7 wt % PAMAM and 10 wt % TGM in pH 7.4 PBS either held at
18.degree. C. for 4 h or held first at 4.degree. C. for 10 min and
then at 37.degree. C. for 3 h.
FIG. 33 shows long-term mass loss due to PAMAM degradation in pH
7.4 PBS of hydrogels with 10 wt % TGM and either 5 or 8 wt %
PAMAM.
FIG. 34 shows viability of rat fibroblasts exposed to soluble
leachables from hydrogels with 7 wt % PAMAM and 10 wt % TGM in 1,
10, and 100.times. dilutions over 2 and 24 h relative to the live
control as determined by fluorescence intensity from Live/Dead
staining with live and dead controls shown.
FIG. 35 shows in vitro leachables cytocompatibility of dual gelling
hydrogels for two TGM wt %. Cell viability for all groups was
greater than 65%. * and # indicate statistical significance
(p<0.05) within and between timepoints, respectively.
FIGS. 36A and 36B show MicroCT images of 20 wt % hydrogels in the 8
mm rat calvarial defect at 4 (FIG. 36A) and 12 (FIG. 36B)
weeks.
FIG. 37 shows the results of microCT scoring for bony bridging and
union for 15 wt % TGM hydrogels and 15 and 20 wt % TGM/DBA
hydrogels at 4 and 12 weeks. Error bars represent standard
deviations for n=6-7. (*) indicates significant change from 4 week
timepoint (p<0.05).
FIG. 38 shows the results of microCT quantification of % bone
volume within the cranial defect at 4 and 12 weeks using an 8 mm
VOI with thresholding gray values (70-255). Error bars represent
standard deviation of n=6-7 hydrogels. (*) and (#) indicate
significant change between timepoints or across groups,
respectively (p<0.05).
FIGS. 39A, 39B, 39C, 39D, 39E, 39F, 39G, 39H, 39I, 39J, 39K, 39L,
39M, 39N, and 39O show representative top (FIGS. A-E) and side
(FIGS. F-J) views of microCT generated three dimensional models
showing mineralization on a binary threshold (70-255) and raw
coronal cross-sections from the center of the defects (FIGS. K-O)
of 15 wt % TGM at 4 (FIGS. A,F,K) and 12 weeks (FIGS. B,C,G,H,L,M)
and 20 wt % TGM hydrogels at 12 weeks (FIGS. D,E,I,J,N,O).
FIGS. 40A, 40B, 40C, 40D, 40E, and 40F show histological staining
of 15 wt % TGM/DBA (left column), 20 wt % TGM/DBA (middle column),
and 15 wt % TGM hydrogels (right column) at 4 (FIGS. A-C) and 12
(FIGS. D-F) weeks. One set of respective images from each timepoint
show von Kossa, hematoxylin and eosin, and Goldner's Trichrome
staining (top to bottom) at 2.times. magnification. Subsets at
bottom demonstrate tissue response at 12 weeks at 20.times.
magnification within the boxed regions. Scale bars for hydrogel
slices represent 1 mm and 8 mm, respectively. Scale bars for high
magnification images represent 100 .mu.m.
FIG. 41 shows a bar graph quantifying the calcium content of 15 wt
% TGM and 15 and 20 wt % TGM/DBA acellular hydrogels (n=4) over 28
days in phosphate buffered saline (PBS), 1.times. simulated body
fluid (SBF), complete osteogenic media without serum (NS), and
complete osteogenic media with serum (S). 15 and 20 wt % TGM/DBA
hydrogels correspond to the left y-axis, and 15 wt % TGM hydrogels
correspond to the right y-axis. (*) refers to significant
difference across timepoints within group and media (p<0.05) and
(#), (&) refers to significant difference of media within group
and timepoint from PBS group, and bars without similar symbols are
significantly different (p<0.05).
FIG. 42A shows the results of histological scoring of scaffold
mineralization across the coronal cross-section in the center of
the defect at 4 and 12 weeks for 15 wt % TGM hydrogels and 15 and
20 wt % TGM/DBA hydrogels. Error bars represent the standard
deviation for n=6-7 samples. (*) and (#) indicates significant
change across groups at same timepoint or from 4 week timepoint,
respectively (p<0.05).
FIG. 42B shows the results of histological scoring of bony bridging
across the coronal cross-section in the center of the defect at 4
and 12 weeks for 15 wt % TGM hydrogels and 15 and 20 wt % TGM/DBA
hydrogels. Error bars represent the standard deviation for n=6-7
samples. (*) and (#) indicates significant change across groups at
same timepoint or from 4 week timepoint, respectively
(p<0.05).
FIG. 43 depicts microCT maximum intensity projections of (A) 15 wt
% TGM 12 wk, (B) 20 wt % TGM/DBA 12 wk, (C) 20 wt % TGM/DBA 4 wk,
and (D) 15 wt % TGM/DBA 4 wk.
While the present disclosure is susceptible to various
modifications and alternative forms, specific example embodiments
have been shown in the figures and are herein described in more
detail. It should be understood, however, that the description of
specific example embodiments is not intended to limit the invention
to the particular forms disclosed, but on the contrary, this
disclosure is to cover all modifications and equivalents as
illustrated, in part, by the appended claims.
DESCRIPTION
The present disclosure generally relates to injectable
compositions. More particularly, the present disclosure relates to
injectable, thermogelling hydrogels and associated methods.
Injectable hydrogel scaffolds can be minimally invasive and can
easily fill complex tissue defects or voids often found in
applications such as craniofacial bone regeneration after trauma,
tumor resection, or birth defects. Thermogelling polymers, such as
poly(N-isopropylacrylamide), which pass through a lower critical
solution temperature (LCST) upon injection into the body, are
promising candidates as scaffold backbones. Concomitant chemical
crosslinking during thermogellation may further enhance the
stability and mechanical properties of such materials; however
these networks may need to be made biodegradable on an appropriate
timescale for tissue regeneration. This may be a major issue for
injectable materials, since many commonly used biodegradable
polymers, such as poly(a-hydroxy esters), may not be water
soluble.
While the majority of effort in tissue engineering research has
been focused on pre-formed implantable scaffolds, there are many
applications that might be better served with injectable, in situ
forming materials capable of co-delivering cells and growth factors
to optimize tissue regeneration. Injectable hydrogel-forming
solutions can be applied via minimally invasive approaches, have
generally high water contents to promote diffusion of nutrients and
cells, and can easily fill complex tissue defects or voids often
found in applications such as craniofacial bone regeneration after
trauma, tumor resection, or birth defects. Various materials
approaches to designing injectable hydrogel constructs have been
previously reviewed.
Thermogelling polymers, which pass through a lower critical
solution temperature (LCST) upon injection into the body, are
promising candidates as scaffold backbones. In particular,
hydrogels based on poly(N-isopropylacrylamide) (pNiPAAm), have been
shown to have versatile and facile application towards the design
of in situ forming hydrogel formulations. Previous efforts have
identified numerous means to alter and control the rate and nature
of the thermally-induced coil-globule phase transition of pNiPAAm
to allow for system optimization, have demonstrated the ability to
successfully deliver encapsulated cells in vitro, and have tested
for efficacy as injectable biomaterials in vivo. However, one major
challenge associated with using thermoresponsive materials for the
design of injectable constructs concerns the tendency of such
polymeric gels to undergo syneresis after initial formation. In
order to explore efficient therapies that ensure complete filling
of tissue defects and promote contact and the exchange of nutrients
and cells with the surrounding native tissue, this limitation must
be addressed.
Recent investigations have focused on the creation of
dual-hardening injectable hydrogels, which combine thermoresponsive
polymers with concomitant in situ crosslinking to stabilize and
strengthen the constructs. While some have utilized click
chemistries to this end, many investigations have created
crosslinking capacity in pNiPAAm hydrogels through polymer
pendant-group modification to introduce reactive double bonds.
However, such systems require generally cytotoxic initiators and/or
catalysts to be included in the injectable formulation to achieve
crosslinking. Macromers of pNiPAAm modified with acrylate and
methacrylate groups have been shown to be cytotoxic as well, thus
necessitating a rapid completion of the crosslinking reaction,
which is generally accomplished through inclusion of catalyst
compounds and/or high initiator concentrations. Despite these
cytotoxicity concerns, pNiPAAm-based dual-hardening hydrogels have
had some success at maintaining the viability of encapsulated
mesenchymal stem cells and promoting construct mineralization. An
additional concern here is the ability of such crosslinks to
degrade on a physiologically relevant timescale, as ester bonds
located proximal to a polymer backbone are quite slow to degrade.
Thus, additional measures must be taken to promote network
degradation, such as synthesis of complex pendant groups or
introduction of degradability into the pNiPAAm backbone.
In contrast to the use of a single macromer to both physically and
chemically gel in situ, two-macromer systems offer a number of
advantages. Key hydrogel properties, including
hydrophilicity/degree of syneresis and the extent of crosslinking
can be altered at the hydrogel formulation stage, as opposed to the
macromer synthesis stage, through combination of a thermogelling
macromer and a hydrophilic crosslinking macromer. By utilizing
cytocompatible and degradable crosslinkers, optimization of the
degradation and thermoresponsive behavior of the hydrogels can be
compartmentalized. This route of construct design also allows for
simplified and cost-effective polymer synthesis and access to a
wide range of crosslinking chemistries. However, material choice is
limited to water-soluble, injectable polymers, which precludes use
of many commonly employed degradable polymers, such as
.alpha.-hydroxy esters.
Polyamidoamines (PAMAMs) are an emerging class of water-soluble and
degradable macromers, thus far used primarily in the design of
dendritic polymers and gene transfection agents. They have reported
biocompatibility and facile synthesis procedures. The nature of the
polyaddition reaction used in PAMAM synthesis, as shown in FIG. 5,
allows for the straightforward production of linear, difunctional
macromers with either amine or acrylamide termini. In addition, the
rate of macromer degradation can be modulated through appropriate
selection of the diamine and bisacrylamide comonomers, and
additional reactive pendant moieties, such as hydroxyl or
carboxylic acid groups, or higher-functionality crosslinkers can be
easily incorporated into the macromer design.
Thus, the present disclosure provides for a novel hydrogel system.
In particular, the hydrogels of the present disclosure are
injectable, thermally responsive, chemically crosslinkable, and
hydrolytically degradable. In certain embodiments, the hydrogel may
serve as an injectable scaffold.
Without wishing to be bound by limitation, it is believed that the
combined use of a physical thermogelling and chemical crosslinking
system allows for enhanced mechanical properties and reduced
syneresis, or the expulsion of liquid from the gel. Previous work
has shown that thermogelling poly(N-isopropylacrylamide)
(PNiPAAm)-based hydrogels crosslinked via methacrylate groups are
cytocompatible and mineralize in vitro. However these
thermosensitive PNiPAAm-based gels undergo significant syneresis
and produce non-soluble degradation products.
The present disclosure provides for a novel system that is capable
of addressing both issues through the inclusion of water-soluble
crosslinkers and a hydrolyzable chemical moiety that impart
non-shrinking and degradative properties to injectable hydrogel
compositions.
The lower critical solution temperature (LCST) is the phase
transition temperature in which a mixture changes from a soluble to
insoluble state. The significance of this temperature in developing
injectable hydrogels of the present disclosure is that a solution
may be injected into the body as a liquid and, upon passing the
LCST, can instantaneously gel and form to the complex space of a
defect or void. FIG. 2 illustrates this concept.
The compositions of the present disclosure generally comprise
thermogelling macromers based on pNiPAAm and hydrophilic and
degradable PAMAM-based macromers. The PAMAM-based macromers serve
as a crosslinker in the compositions of the present disclosure. The
compositions of the present disclosure may allow for
dual-hardening, injectable hydrogels. One advantage of the hydrogel
compositions of the present disclosure is that the rate of
degradation of the hydrogels may be easily tuned through alteration
of the PAMAM backbone. This is especially beneficial for tissue
engineering application.
In certain embodiments, the thermogelling macromers may have the
following structure:
##STR00001##
wherein x, y, z, and m are set forth as shown in FIG. 4, and
wherein x is in the range of from about 0 to about 10 mol %, z is
in the range of from about 0 to about 10 mol %, m is in the range
of from about 2.5 mol % to about 15 mol %, and y is in the range of
from about 65 mol % to about 97.5 mol %. n may be greater than or
equal to 1. The molecular weight of the thermogelling macromers may
vary depending on the intended application of the macromers. In
certain embodiments, the number average molecular weight can vary
from about 2000 to about 100,000 Da. In certain embodiments, when
the composition may be used for a mammal, the number average
molecular weight may be less than 32,000 Da. The order of the
components of the thermogelling macromer are not limited to the
order illustrated above.
In certain embodiments, the pNiPAAm-based macromers may be
co-polymerized with other groups that may aid in modulating LCST.
For example, the pNiPAAm-based macromers may be copolymerized and
modified with pendant lactone rings. Copolymerizing the
pNiPAAm-based macromers with pendant lactone rings may enable
hydrolysis-dependent degradation via LCST modulation. Incorporation
of the ringed monomer, for example, dimethyl-.gamma.-butryolactone
(DBA), may provide hydrolysis-dependent LCST-modulating abilities
over time for the creation of soluble byproducts. FIG. 3
illustrates this basic concept. For example, DBA comonomer may
modulate LCST through hydrolysis of lactone rings and DBA comonomer
may decrease LCST. The presence of hydroxyl groups increases the
hydrophilicity of thermogelling macromers so that the polymer may
revert back to soluble macromer units. FIG. 3A illustrates
hydrolysis of DBA. FIGS. 3B and 3C show the LCST modulation effect
that DBA may have on a polymer with varying DBA content. Other LCST
modulating groups that may be copolymerized with pNiPAAm-based
macromers include but are not limited to acrylic acid (AA),
glycidyl methacrylate (GMA), acrylamide (AAm) and hydroxyethyl
acrylate (HEA). For example, AA and AAm comonomers may increase
LCST while GMA may decrease LCST. The degree to which the LCST
modulation is desired will intend on the intended application of
the composition of the present disclosure.
In one embodiment, thermogelling macromers may be p(NiPAAm-co-GMA).
In other embodiments, the thermogelling macromers may be
p(NiPAAm-co-GMA-co-DBA-co-AA). In other embodiments, the
thermogelling macromers may be p(NiPAAm-co-AAm-co-HEA)). In other
embodiments, the thermogelling macromers may be
p(NiPAAm-co-GMA-co-AAm).
In one embodiment, the PAMAM-based macromers may be an
epoxy-reactive polyamidoamine diamine crosslinkers. Diamine
functionalized polyamidoamine crosslinkers may increase the
equilibrium degree of swelling through epoxy-based reactions with
glycidyl methacrylate moieties of the base polymer to create
non-shrinking injectable hydrogels. In another embodiment, the
PAMAM-based macromers may be functionalized with bisacrylamide.
PAMAM-based macromers used in conjunction with the compositions of
the present disclosure have certain advantages. They are easily
synthesized, allowing for flexibility. In certain embodiments, they
may be the polycondensation product of bisacrylamide and diamine
monomers. They are biodegradable through amide bonds, water
soluble, and have reported biocompatibility.
The compositions of the present disclosure may also be used to
deliver cells or other biochemical agents. In certain embodiments,
PNiPAAm-based macromers may be combined with cells. In certain
embodiments, the cells may be multipotent mesenchymal stem cells.
The mesenchymal stem cells may be subsequently encapsulated and
evenly dispersed within the polymer network upon thermogelation
above the LCST. In certain embodiments, the hydrogel compositions
of the present disclosure may provide a means by which localized
and minimally invasive cell delivery may be achieved within a
mammal. In certain other embodiments, the hydrogel compositions of
the present disclosure may be used to deliver growth factors or a
combination of growth factors and cells to a specific location
within a mammal.
The compositions of the present disclosure may provide certain
advantages. Network formation through the epoxy crosslinking
reaction may be rapid and facile. In certain embodiments, the
reaction may reach completion in less than three hours after an
initial thermogellation time of 2-3 seconds. Additionally, the
often problematic tendency of thermogelling systems to undergo
significant post-formation gel syneresis may be mitigated through
the combination of increased hydrogel hydrophilicity and gel
hardening through concomitant chemical crosslinking during and
after initial thermogellation.
Such in situ dual-hardening, dimensionally stable, defect-filling,
and degradable hydrogels with high gel water content are attractive
substrates for tissue engineering and cellular delivery
applications. In particular, the use of water-soluble and
degradable polyamidoamine macromers offers tremendous synthetic
flexibility and control over subsequent gel properties. The
hydrogel compositions of the present disclosure also allow for the
ability to tune hydrogel hydrophilicity, degree of post-formation
swelling or syneresis, degradation timescale, degree of
crosslinking, and potential introduction of additional pendant
functional moieties with appropriate selection of starting
comonomers.
The compositions of the present disclosure solidify or gel thorough
a dual-gelation, physical and chemical mechanism upon preparation
and elevation of temperature. In certain embodiments, the
temperature may be approximately 37.degree. C. In certain
embodiments, the compositions may solidify in situ. In certain
embodiments, the increased hydrogel hydrophilicity due to
increasing PAMAM-based macromer incorporation into the polymer
network mitigates the often problematic tendency of thermogelling
materials to undergo significant syneresis after hydrogel
formation. In certain embodiments, an epoxy-based crosslinking
reaction allows for rapid and facile incorporation of the
PAMAM-based macromers into the polymer network during and after
thermogellation. In certain embodiments, when thermogelling
macromers contains GMA, the compositions may solidify and form via
epoxy reactive crosslinking between GMA moieties on the
thermogelling macromer and PAMAM units. (FIG. 6).
In certain embodiments, the thermogelling macromers based on
pNiPAAm may be synthesized by free radical polymerization. In
certain embodiments, the PAMAM-based macromers may be synthesized
by polymerization of methylene bisacrylamide and piperazine.
To form the hydrogel compositions of the present disclosure,
pNiPAAm-based macromers and PAMAM-based macromers are combined. The
pNiPAAm-based macromers and PAMAM-based macromers may be combined
in any ratio depending on the intended application and desired
properties of the hydrogel. In certain embodiments, the
pNiPAAm-based macromers may be present in the hydrogel composition
in a range of from about 5 wt % to about 25 wt %. In certain
embodiments, the pNiPAAm-based macromers may be present in the
hydrogel composition in the range of from about 10 wt % to about 20
wt %. In certain embodiments, pNiPAAm-based macromers may be
present in the hydrogel composition in the range of from about 15%
to about 20%.
The PAMAM-based macromers may be present in the hydrogel
composition of the present disclosure in an amount dependent on the
amount of epoxy ring functional groups on the pNiPAAm-based
macromer backbone. In certain embodiments, the PAMAM-based
macromers may be present in the hydrogel composition based on the
functionality ratio (molar ratio of PAMAM amine functional groups
to epoxy ring functional groups.) In certain embodiments, the the
PAMAM-based macromers may be present in the hydrogel composition in
an amount in the range of from about 1.3 wt % to about 31.2 wt % of
the hydrogel composition. However, the amount of PAMAM-based
macromers and pNiPAAm-based macromers may be varied depending on
the intended application of the hydrogel compositions.
The resulting macromer solution may then be exposed to an elevated
temperature to allow for the solution to solidify due to
thermogellation. In certain embodiments, the temperature may be
about 37.degree. C. The solidified hydrogels are stabilized and
further hardened through epoxy-based chemical crosslinking, which
creates a degradable polymer network structure.
In one embodiment, the present disclosure provides an injectable,
thermogelling tissue engineering scaffold that comprises
polyamidoamine. In another embodiment, the present disclosure
provides an injectable PNiPAAm-based scaffold with tunable
LCST.
The hydrogel compositions of the present disclosure may be used as
an injectable scaffold to treat a defect within a mammal. In
certain embodiments, the defect may be a craniofacial defect.
Injectable, in situ forming materials capable of delivering both
cells and growth factors are viable alternative treatments for the
regeneration of bone tissue in complex craniofacial defects. Tissue
engineering strategies involving injectable, in situ forming
hydrogel scaffolds capable of mesenchymal stem cell (MSC) delivery
show promise for regenerating complex craniofacial defects.
Hydrogels based on Poly(N-isopropylacrylamide) (PNiPAAm) are
particularly attractive since MSCs can be easily mixed with the
polymer solution at room temperature, and subsequently be
encapsulated and evenly dispersed within the insoluble network upon
thermogelation above the lower critical solution temperature
(LCST).
Delivery of multipotent mesenchymal stem cells (MSCs) within in
situ hardening hydrogel networks may be a viable alternative for
bone regeneration in defects of any shape. It has been has shown
that MSC encapsulation in a thermally and chemically crosslinked
poly(N-isopropylacrylamide) (PNiPAAm)-based hydrogel consisting of
PNiPAAm, pentaerythritol diacrylate monostearate (PEDAS),
acrylamide (AAm) and 2-hydroxyethyl acrylate (HEA) may induce
osteogenic differentiation and mineralization in vitro. Klouda et
al. demonstrated these hydrogels enabled osteogenic differentiation
of encapsulated MSCs in vitro.
The hydrogel compositions of the present disclosure are generally
biodegradable and produce non-toxic degradation products that are
soluble. The rate of degradation of the solidified hydrogel
likewise should not impede neotissue formation when the hydrogel
composition is used as a scaffold in a mammal. Moreover, the
hydrogel compositions of the present disclosure should be
biocompatible.
To facilitate a better understanding of the present disclosure, the
following examples of certain aspects of some embodiments are
given. In no way should the following examples be read to limit, or
define, the entire scope of the disclosure.
EXAMPLES
Example 1
Injectable, PNiPAAm-Based Scaffolds with Tunable LCST for
Craniofacial Bone Regeneration
Some objectives of this example were to synthesize a thermogelling
PNiPAAm-based polymer incorporating DBA, acrylic acid (AA) and
glycidyl methacrylate (GMA) monomer units, combine it with a
polyamidoamine (PAMAM), epoxy-reactive crosslinker to reduce
syneresis, and investigate the swelling, degradation and
compressive mechanical properties of these injectable
scaffolds.
Materials and Methods:
Thermogelling macromers (TGMs) were prepared with PNiPAAm, GMA, AA
and DBA, via free radical polymerization. Different compositions of
PAMAM crosslinker were created using a simple polymerization
following established protocols. Characterization of the TGMs and
percent incorporation of AA were found using proton nuclear
magnetic resonance spectroscopy (.sup.1H NMR) and titration,
respectively. The effect of pH and percent DBA incorporation on the
initial LCST was measured in a one month degradation study at
37.degree. C. using differential scanning calorimetry (DSC) and 1H
NMR. A factorial design study was performed to determine the
effects of high and low TGM weight percent, DBA incorporation,
PAMAM crosslinker length and crosslinker density on gelation and
equilibrium swelling (FIG. 7). Several conditions were selected for
further mechanical testing utilizing a thermomechanical analyzer
and degradation studies.
Results and Discussion:
Gels using thermoresponsive, chemically crosslinked macromers are
illustrated in FIG. 8. TGMs with GMA crosslinked with PAMAM
thermogelled and crosslinked rapidly without syneresis at
37.degree. C. Swelling studies demonstrated that gels with higher
crosslinking densities and lengths exhibited greater swelling than
other groups. (FIG. 7). Degradation studies without the crosslinker
showed that at neutral and basic pH, hydrolysis of the DBA lactone
ring was enhanced, raising the initial peak LCST from
26.0+0.6.degree. C. to about 31.degree. C. in 28 days (as shown in
FIG. 9), which was mirrored in further gelation studies. The loss
of thermogelling activity (indicated by the color change in FIGS.
8b and 8c) was caused by an increase in the LCST above 37.degree.
C. due to the basic PAMAM crosslinker, which rapidly catalyzed the
degradation of DBA, leaving a fully crosslinked, but no longer
thermogelled, hydrogel after 30 minutes.
Conclusions:
The presence of dual thermal and chemical hardening mechanisms
allowed for instantaneous scaffold formation upon injection and
reduced subsequent hydrogel syneresis, which is beneficial for the
3D incorporation and proliferation of cells. A promising system was
developed for injectable craniofacial tissue regenerating
therapies. It was also determined that dimethyl-y-butryolactone
successfully modulates LCST, PAMAM-GMA epoxy crosslinking with
PNiPAAm can create thermogelling and chemically crosslinking gels,
and hydrogel properties can be tuned with different parameters.
Example 2
Injectable, Thermogelling Tissue Engineering Scaffolds Utilizing
Polyamidoamines
Materials and Methods:
PNiPAAm was copolymerized firstly with 2-hydroxyethyl acrylate
(HEA) and acrylamide and secondly with glycidyl methacrylate (GMA)
following previously established synthesis and purification
protocols. Polymer composition was determined by .sup.1H NMR;
molecular weight by GPC; LCST by differential scanning calorimetry
(DSC) and rheology. Polyamidoamine crosslinkers were created from
piperazine and methylene bisacrylamide following previous synthesis
procedures with both diamine (PAMAM) and diacrylamide (P-BA)
products. Molecular weight and product distribution was determined
by MicroTOF mass spectrometry. Polymers incorporating HEA were
chemically modified with methacryloyl chloride according to
established protocols to introduce reactive double bonds, combined
with P-BA crosslinkers, and scaffold formation was performed as
previously reported. Polymers containing GMA were combined with
PAMAM crosslinkers and injected into moulds with varying setting
times after mixing. Factorial designs were performed on both
systems to evaluate the effects of PAMAM molecular weight,
PAMAM/crosslinking density, and wt % of thermogelling polymer used
for injection. Select systems were further evaluated with
mechanical testing and degradation studies.
Results and Discussion:
A significant concern often encountered in systems seeking to
utilize a polymer's LCST for in situ scaffold formation is the
subsequent syneresis of the gel after initial formation. A
previously developed system was optimized to reduce and eliminate
observed scaffold syneresis and enhance network degradation on an
appropriate timescale through the introduction of P-BA as a
crosslinking molecule to form gels such as the one seen in FIG.
10.
Separately, a system was designed to leverage the benefits of
polyamidoamines to avoid the use of free radical crosslinking
pathways in favor of an epoxy-based crosslinking mechanism. PAMAM
crosslinkers were combined with p(NiPAAm-co-GMA) to produce
thermogelling and rapidly crosslinking gels. The extent of
syneresis or swelling after initial formation was tunable through
appropriate selection of system variables. FIG. 11 shows weight
swelling ratios at formation and equilibrium for 10 and 20 wt %
thermogelling polymer solutions in PBS (pH=7.4) with a 1.8 kDa
PAMAM crosslinker with three values of crosslinker incorporation.
FIG. 12 shows accelerated (pH=10) degradation profiles of two
formulations with varying PAMAM molecular weight (1.8 and 2.6
kDa).
Conclusions:
Polyamidoamines were shown to be a versatile and promising material
for tissue engineering, particularly in injectable applications.
The systems designed were highly tunable, from the degree of
syneresis or swelling of scaffolds post-formation to the timescale
of degradation of the polymer network.
Example 3
The objectives of this example were: (1) develop novel, injectable,
dual-hardening hydrogel scaffolds for craniofacial tissue
regeneration, (2) synthesize and characterize
P(N-isopropylacrylamide)-based thermogelling macromers and
polyamidoamines, (3) determine efficacy and kinetics of in situ
hydrogel formation, and (4) investigate effects of macromer
properties and formulation parameters on hydrogel swelling,
degradation, and moduli. General approaches taken include studying
(1) injectable, in situ hardening materials, (2) the
thermogellation mechanism, (3) in situ crosslinking (the desire to
avoid syneresis and complete crosslinking over 2 hours or less),
and (4) biodegradability (non-toxic degradation products, soluble
degradation products, and appropriate degradation rates).
FIG. 13 illustrates a polyamidoamine. The polyamidoamine may be a
polycondensation product of bisacrylamide and diamine monomers. The
benefits of using these compounds are: (1) the macromers are water
soluble, (2) they are biodegradable through amide bonds, (3) they
are reported to be compatible, and (4) they have synthetic
flexibility to tune degradation rate.
FIG. 14A illustrates one synthesis scheme for a thermogelling
macromer (TGM). FIG. 14B also shows .sup.1H NMR spectrum for the
TGM. TGM characterization results indicated that the TGM had 7.44
mol % GMA. M.sub.n was equal to 8.86 kDa as determined by GPC.
Finally, the LCSTs of the TGM were determined by DSC. The onset and
peak LCST at 10 wt % in PBS pH 7.4 were 22.2.+-.0.7.degree. C. and
30.9.+-.0.4.degree. C., respectively.
FIG. 15 illustrates the synthesis of a polyamidoamine (PAMAM) and
its characterization.
FIG. 16 illustrates a TGM/PAMAM gelation factorial study. In
general, PAMAM and p(NiPAAm-co-GMA) were mixed at 4.degree. C. The
mixture was then injected in molds at 37.degree. C. After 24 hours,
the w.sub.f (formation) was measured and the mixture was
transferred to PBS pH=7.4. After 48 hours, w.sub.eq (equilibrium)
and lyophilized w.sub.d (dry) were measured. DSC shows crosslinking
reaction was completed before two hours (10 wt % TGM and 1:1 Func.
Ratio).
FIG. 17 illustrates gelation and equilibrium swelling. The results
indicate the production of non-shrinkable hydrogels. The results
also demonstrated tunable post-gelation hydrogel expansion. Lower
gelation swelling at 20 wt % occurred due to more polymer per
volume. Equilibrium swelling was consistent for 10 wt % and 20 wt %
gels.
FIG. 18 illustrates factorial study results. The results indicated
that longer PAMAMs increased swelling, longer preparation time
increased swelling, and the degree of crosslinking had little
effect, while the hydrophilicity of the hydrogel dominated. Higher
polymer content improved compressive strength, while Young's
Modulus decreased by approximately 50$ for 0.5:1 func. ratio.
FIGS. 19 and 20 illustrates PAMAM degradation results. Higher
molecular weight PAMAM exhibited slower rate of mass loss. This
indicated that it is likely that hydrophilicity of hydrogel
controls early syneresis. Increased polymer content slowed rate of
gel syneresis/mass loss. Hydrogel degradation at pH 7.4, 37.degree.
C. was completed for all samples studied by 12 weeks. There was a
moderate (10-25%) mass loss until 9+ week.
FIG. 21 illustrates a PAMA-BA system. FIG. 22 illustrates the
results of degree of modification. FIG. 23 illustrates one example
of a PAMAM-BA.
The conclusions of this example were: (1) two promising, novel
hydrogel systems for biomaterials application, particularly
injectable scaffolds, were synthesized, (2) hydrogels were shown to
avoid often problematic issues of syneresis, (3) hydrogels were
shown to degrade in an appropriate timescale for tissue
regeneration applications (within 12 weeks), and (4) hydrogel
formulation conditions were investigated and gel syneresis,
mechanical properties and degradation timescales were highly
tunable, though intertwined.
Example 4
Thermogelling Macromer (TGM) Synthesis
N-isopropylacrylamide (NiPAAm), glycidyl methacrylate (GMA),
2,2'-azobis(2-methypropionitrile) (azobisisobutyronitrile, AIBN),
N,N'-methylenebisacrylamide (MBA), piperazine (PiP), and 0.1N
sodium hydroxide solution were purchased from Sigma-Aldrich (Sigma,
St. Louis, Mo.) and used as received. The solvents; tetrahydrofuran
(THF), dimethylformamide (DMF), diethyl ether, and acetone in
analytical grade, and water, acetonitrile, chloroform, and methanol
in HPLC-grade; were obtained from VWR (Radnor, Pa.) and used as
received. Phosphate-buffered saline (PBS) solution was mixed from
powder (pH 7.4, Gibco Life, Grand Island, N.Y.), and ultrapure
water was obtained from a Millipore Super-Q water system
(Millipore, Billerica, Mass.).
The thermogelling macromer p(NiPAAm-co-GMA) was synthesized by free
radical polymerization as shown below. Ten grams of the comonomers,
NiPAAm and GMA at 92.5 and 7.5 mol %, respectively, were dissolved
in 100 mL of DMF and polymerized at 65.degree. C. under nitrogen
atmosphere. AIBN was added as a free radical initiator at 0.7 mol %
of the total monomer content, and the reaction mixture was
continuously stirred for 20 h. The product was concentrated by
rotary evaporation, dissolved in THF, and twice precipitated in at
least 10-times excess of cold diethyl ether to effectively remove
the unreacted monomers and low molecular weight oligomers. The
final filtrate was dried under vacuum at ambient temperature to
yield a fine white powder. In addition, a similar procedure was
followed to create a pNiPAAm homopolymer.
##STR00002##
Homopolymers of pNiPAAm and copolymers with GMA were successfully
synthesized to yield thermoresponsive macromers with functional
pendant epoxide moieties. The number average molecular weight,
M.sub.n, and polydispersity index (PDI) of the synthesized pNiPAAm
macromer were 8.9.+-.0.1 kDa and 2.72.+-.0.08, respectively, as
determined by GPC (data not shown). Similarly, values of
M.sub.n=9.2.+-.0.6 kDa and PDI=3.12.+-.0.04 were found for
p(NiPAAm-co-GMA).
Copolymer composition was further evaluated by .sup.1H NMR. FIG. 24
shows the .sup.1H NMR spectra of p(NiPAAm-co-GMA). The five
individual pendant proton signals (3d and 2e) on GMA were found
between 2.5 and 4.8 ppm, along with the isopropyl proton from
NiPAAm (1a, 3.8-3.9). The six NiPAAm methyl protons (6b) were found
from 0.9-1.3, and the polymer backbone protons (3c, 3f) were
located from 1.3-2.3 ppm. The proton locations were confirmed by
comparison with spectra of the NiPAAm and GMA monomers and the
pNiPAAm homopolymer (data not shown). The varied peak locations
allowed for five independent calculations of the GMA content in the
copolymer through the relative intensities of the GMA and NiPAAm
protons giving an average value of 7.44.+-.1.12 mol %, which
compares well with the theoretical GMA content (7.5 mol %).
Finally, the LCSTs of the homopolymer and copolymer were determined
by DSC (FIG. 25). The onset and peak LCST for pNiPAAm were
29.5.+-.0.2.degree. C. and 33.3.+-.0.2.degree. C., respectively,
which compare well with the commonly reported LCST of 32.degree. C.
Copolymerization with GMA reduced the LCST to an onset of
22.2.+-.0.7.degree. C. and a peak of 30.9.+-.0.4.degree. C. In
addition, p(NiPAAm-co-GMA) displayed lower overall transition
energy and a broader temperature distribution compared with the
homopolymer, which can be attributed to lower overall NiPAAm
content and greater chain content variation in the copolymer,
respectively. The onset LCSTs as determined by DSC correlated very
well to macroscopic observation of the temperatures at which
thermogellation occurred (data not shown).
Example 5
Polyamidoamine Synthesis
The polyamidoamine (PAMAM) crosslinking macromers were synthesized
by polyaddition of PiP with MBA (as shown below). 10.83 g of the
comonomers were dissolved in 30 mL ultrapure water with a
stoichiometric excess of PiP (r=[MBA]/[PiP]=0.846), stirred
continuously under nitrogen atmosphere at 30.degree. C., and
allowed to react for 48 h according to published procedures
(Ferruti 2002 and Dey 2005). The obtained viscous mixture was
directly precipitated in 100 mL acetone, filtered, and dried under
vacuum at ambient temperature to yield a fine powder.
##STR00003##
An epoxy-reactive polyamidoamine diamine crosslinker was
successfully synthesized by adapting reported protocols. (Ferruti
2002 and Dey 2005). FIG. 26 shows the .sup.1H NMR spectrum of the
final purified and dried product of the PAMAM synthesis. The
expected molecular structure of the diamine product is shown with
protons and their corresponding peak locations are identified. The
peak locations were confirmed by analysis of the MBA and PiP
comonomers (data not shown). A crude estimate of the number average
degree of polymerization can be calculated based on functional
group quantification and assuming 100% conversion of the limiting
reactant (MBA) as shown in FIG. 26. Application of these equations
yielded r=0.840 and X.sub.n=11.52 (M.sub.n=1350). These values
compared favorably with the feed stoichiometric ratio (r=0.846) and
expected molecular weight (X.sub.n=11.99, M.sub.n=1400 D), assuming
100% conversion of MBA.
This molecular weight calculation assumed that all of the
functional groups present were part of chains of the product
diamine. However, since the product is susceptible to hydrolytic
degradation and the synthesis was run in water, further analysis of
the actual product compositions was necessary. To this end,
time-of-flight mass spectroscopy analysis with electrospray
ionization (microTOF mass spectroscopy) was utilized to identify
and quantify the distribution of chain species present. FIG. 27
displays the resulting raw data of signal intensity versus absolute
mass with major peaks and associated charge states identified.
Charge states of peak families were determined through measurement
of the separation of isomeric peaks, as shown in the insert.
MicroTOF mass spectroscopy led to the successful identification,
based on expected degradation products and their subsequent
potential reactions, of all major species present in the final
PAMAM product, as summarized in Table 1. Each compound identified
appeared at numerous molecular weights and varying charge states.
In addition to proton charges, sodium and occasional potassium
adducts were also observed. Thus, the peaks were deconvoluted,
summed, and assigned structures with matching absolute molecular
weight values: (a) is the desired diamine PAMAM product; (b) is the
heteroend version of the desired product; (c) and (g) are simple
degradation products; and the remainder are secondary products of
the reaction of either the primary amine groups ((e) and (f)) or
the carboxylic acid groups ((d) and (h)) resulting from hydrolytic
degradation of the PAMAM amide linkages.
TABLE-US-00001 TABLE 1 Molecular structure of primary, degradation,
and secondary products of the PAMAM synthesis, as determined by
microTOF mass spectroscopy, and their mol % of the total species
present. Mol Structure % ##STR00004## (a) 88.5 ##STR00005## (b) 3.7
##STR00006## (c ) 2.7 ##STR00007## (d) 0.5 ##STR00008## (e) 1.3
##STR00009## (f) 1.4 ##STR00010## (g) 1.6 ##STR00011## (h) 0.2
It should also be mentioned that each product structure shown is
the representative or simplest form of that particular product,
with the degradation and secondary reaction products shown at the
end of a PAMAM chain. However, further addition of repeat units is
often possible beyond the end groups. Thus, each degradation and
secondary reaction product shown represents a class of polymer
species having the same molecular weight and functional end groups.
Also shown in Table 1 are the mol percents of each species
identified, as determined from the relative microTOF peak
intensities, and a further product analysis is presented in Table
2.
TABLE-US-00002 TABLE 2 PAMAM synthesis product distribution (mol %
and wt %). Synthesis Product Mol % Wt % Desired Diamine 88.5% 95.2%
Heteroend Product 3.7% 1.3% Degradation Products 7.8% 3.5%
Mono-amine 4.7% 2.0% Other 3.1% 1.5%
The desired PAMAM diamine comprises 88.5 mol % and 95.2 wt % of the
total reaction product, with the difference a result of the
inherently smaller nature of species originating from degradation
pathways. Including the side-products and degradation-derived
products of the synthesis scheme utilized, 98.5 wt % of the
resultant powder will participate in epoxy reactions with the TGM
macromer, with 95.2 wt % having the potential to create crosslinks
and 3.3 wt % simply the potential to form hydrophilic grafts. The
remaining 1.5 wt % is comprised of species with non-amine
functional ends, namely carboxylic and acrylamide functional ends
(structures shown in Table 1). It should be mentioned, however,
that the accuracy of such compositional quantifications could be
limited by the nature of the mass spectroscopy analysis. Namely, we
are assuming that all species are equally ionizable, evenly
electrosprayed, and the detection signal is linear over more than
two orders of magnitude.
Based on initial experimental observation, it was hypothesized that
the polyamidoamine reaction likely was completed much earlier than
the 48 h reaction time period, which was chosen based on published
reports. (Dey 2005). The hope was that optimizing the reaction
timescale would result in fewer degradation and secondary reaction
products. To this end, the reaction kinetics were investigated by
sampling the reaction mixture over time and performing microTOF
mass spectroscopy. The results of the transient reaction products
study are summarized in FIG. 28, where M.sub.n and the total
desired product diamine weight fraction (of the total species
present in the reaction media) are shown over the course of the
reaction (at 30 min, 1, 2, 4, 12, 24, and 48 h).
As hypothesized, the reaction primarily occurs and has nearly
completed within the first 4 h, however no further change in the
content of degradation and secondary reaction products was seen
beyond the completion of the primary reaction. The absence of such
a trend in retrospect can be attributed to the changing nature of
the reaction media as functional groups are consumed. Namely, the
rapid consumption of amine groups during the polyaddition reaction
led to a significant decrease in the availability of base groups
with associated decreases in the pH of the reaction solution (FIG.
28 insert). Since the hydrolytic degradation is base catalyzed, the
vast majority of degradation could be anticipated to occur in the
very early stages of the synthesis procedure.
This hypothesis is borne out in preliminary PAMAM and hydrogel
degradation studies, where the rate of degradation is more than an
order of magnitude faster at pH 10 compared to pH 7.4. The .about.2
kDa M.sub.n (with a PDI of .about.1.5) is higher than that found
from .sup.1H NMR analysis and expected from theory, but represents
only the desired diamine distribution. When all species were
included in the calculation, the M.sub.n was .about.1.8 kDa.
Finally, alternate solvents were investigated for the synthesis of
the PAMAM diamine that might result in less product degradation,
however due to component solubilities and the need for a protic
solvent, an alternative was not identified.
Example 6
Methods for Proton Nuclear Magnetic Resonance Spectroscopy (.sup.1H
NMR)
.sup.1H NMR spectra were obtained using a 400 MHz spectrometer
(Bruker, Switzerland). Sample materials were dissolved in D.sub.2O
(typical concentration: 20 mg/mL) that contained 0.75 wt %
3-(trimethylsilyl)propionic-2,2,3,3-d.sub.4 acid, sodium salt (TSP)
as internal shift reference (Sigma-Aldrich, St. Louis, Mo.). All
post-acquisition data processing was performed with the MestRe-C
NMR software package (Mestrelab Research S.L., Spain). The free
induction decay (FID) was Fourier transformed, manually phased,
referenced using the TSP signal, baseline corrected, and
integrated.
Example 7
Gel Permeation Chromatography (GPC)
Molecular weight distributions of the p(NIPAAm-co-GMA) and pNiPAAm
polymers were determined by GPC. A GPC system consisting of an HPLC
pump (Waters, model 510, Milford, Mass.), an autosampler/injector
(Waters, model 717), and a differential refractometer (Waters,
model 410) equipped with a series of analytical columns (Styragel
guard column 20 mm, 4.6.times.30 mm; Styragel HR3, 5 mm,
4.6.times.300 mm; Styragel HR1 column, 5 mm, 4.6.times.300 mm (all
Waters)) was used with degassed chloroform as the eluent at a flow
rate of 0.3 mL/min. Samples were prepared in chloroform at a
concentration of 25 mg/mL and filtered prior to analysis. Macromer
number average molecular weight (M.sub.n) and polydispersity index
(PDI) were determined in triplicate relative to polystyrene
standards.
Example 8
MicroTOF Mass Spectroscopy
Molecular weight distributions of the synthesized PAMAM crosslinker
were analyzed using time-of-flight mass spectroscopy with
positive-mode electrospray ionization on a Bruker microTOF ESI
spectrometer (Bruker Daltonics, Billerica, Mass.) equipped with a
1200 series HPLC (Agilent Technologies, Santa Clara, Calif.) to
deliver the mobile phase (50:50 HPLC-grade water and methanol).
During the PAMAM synthesis, 100 .mu.L samples of the reaction
mixture were collected at 30 min and 1, 2, 4, 12, 24, and 48 h,
diluted with 550 .mu.L of a 50:50 mixture of HPLC-grade water and
acetonitrile with 0.1% formic acid added, and a 2 .mu.L flow
injection was delivered to the electrospray chamber. After data
acquisition, all peaks (including degradation and secondary
reaction products) were identified using microTOF Control software
(Bruker), corrected for charge state (generally with H.sup.+ or
Na.sup.+ and rarely K.sup.+ ions), and quantified for calculation
of M.sub.n, M.sub.w, and PDI.
Example 9
Differential Scanning calorimetry (DSC)
The LCSTs of the thermogelling macromers were determined by DSC. 14
.mu.L of 10 wt % polymer in pH 7.4 PBS solutions were pipetted into
aluminum volatile sample pans (TA Instruments, Newcastle, Del.) and
capped/crimped. Thermograms were recorded in triplicate on a TA
Instruments DSC 2920 equipped with a refrigerated cooling system
against an empty sealed pan as reference. In a typical run, the
oven was equilibrated at either 5.degree. C. or -5.degree. C. for
10 min and then heated to 80.degree. C. at a heating rate of
5.degree. C./min. The LCST was determined both as the onset and
peak temperature of the endothermic peak in the thermogram using
the Universal Analysis 2000 software provided with the DSC system.
In addition, the progress of the hydrogel formation/crosslinking
reaction was monitored by DSC. 14 .mu.L of the hydrogel solution
(10 wt % TGM and 7 wt % PAMAM in pH 7.4 PBS) was pipetted into an
aluminum pan, capped/crimped, placed into the DSC and equilibrated
at 37.degree. C., and monitored until the completion of the
reaction.
Example 10
Rheological Characterization
A thermostated, oscillating rheometer (Rheolyst AR1000, TA
Instruments, New Castle, Del.) equipped with a 6 cm steel cone (1
degree) with a gap size of 26 .mu.m was used to evaluate the
elastic response of the hydrogels. Injectable hydrogel formulations
containing 7 and 10 wt % (w/v) of PAMAM and TGM, respectively, in
pH 7.4 PBS were pipetted onto the rheometer, and the dynamic
viscoelastic properties of the solutions, namely, the dynamic shear
storage (G') and loss (G'') moduli, complex viscosity (|.eta.*|),
and loss angle (.delta.), were recorded using the TA Rheology
Advantage software (TA Instruments) as the solution was either
maintained at a temperature of 18.degree. C. for 4 h or maintained
at 4.degree. C. for 10 minutes followed by rapid elevation to and
maintenance at 37.degree. C. for 3 h.
Example 11
Hydrogel Formation and Degradation
Individual solutions of the TGM and PAMAM macromers were prepared
at twice the desired final solution concentrations, and 250 .mu.L
of each TGM and PAMAM solution combination were combined in a glass
vial at 4.degree. C. and mixed gently for .about.30 s. The
injectable solutions were then immediately immersed in a 37.degree.
C. water bath and allowed 24 h to reach equilibrium. After
equilibrium swelling analysis, degradation of the PAMAM networks at
37.degree. C. was evaluated in both pH 7.4 PBS and 0.1N sodium
hydroxide solution.
The TGM and PAMAM macromers were combined as synthesized in pH 7.4
PBS to create injectable, in situ dual-hardening solutions. Upon
macromer mixing and increasing the temperature to 37.degree. C.,
the solutions will rapidly solidify due to the thermogellation
mechanism. Subsequently, the hydrogels will be stabilized and
further hardened through epoxy-based chemical crosslinking to
create a degradable polymer network structure (FIG. 29).
FIG. 30 illustrates the effect of progressive incorporation of the
PAMAM crosslinker into 10 wt % TGM injectable solutions, with the
equilibrium volume swelling ratio relative to the 0 wt % PAMAM
hydrogel, Q.sub.r, shown. As hypothesized, the increased hydrogel
hydrophilicity mitigated the often problematic tendency of
thermogelling solutions to undergo significant post-formation
syneresis, while maintaining the ability of the TGM to undergo
thermogellation at 37.degree. C. It should also be noted that the
theoretical maximum degree of crosslinking for this particular
combination of TGM and PAMAM macromers is .about.7 wt % PAMAM for
10 wt % TGM, as determined from quantification of reactive amine
and epoxy functionalities per unit mass for each macromer. Thus,
beyond 7 wt % PAMAM, the fraction of PAMAM molecules forming
intermolecular crosslinks or intramolecular loops will decrease as
some are replaced by hydrophilic branches.
The kinetics of the epoxy crosslinking reaction were first
monitored thermally by DSC for injectable hydrogel solutions with
10 and 7 wt % of TGM and PAMAM, respectively. As can be seen in
FIG. 31, the crosslinking reaction was completed within 110
minutes. Such rapid in situ hardening of the thermogelled hydrogel
constructs enhances gel stability while minimizing exposure of
encapsulated cells and surrounding tissue to reactive species.
Further evidence of the progress of the crosslinking reaction was
seen in rheology traces. FIG. 32 shows the rheological response of
the hydrogels to crosslinking both in the presence (4/37.degree.
C.) and absence (18.degree. C.) of physical thermogellation. Thus,
the dual, thermally and physically, gelling nature of the hydrogel
system is clearly illustrated.
Network formation in the absence of thermogellation at 18.degree.
C. resulted in hydrogels that reached the gel point in .about.100
min and had ultimate shear storage moduli, G', on the order of 1
kPa. Network formation in the presence of thermogellation at
37.degree. C. reached the gel point at .about.60 min and showed the
relative contributions of the physical and thermal gelation
mechanisms, with an ultimate shear storage modulus, G', on the
order of 90 kPa. The timescale of the crosslinking reaction as
shown by rheology was somewhat lengthier than that shown by DSC
(.about.180 and .about.110 min, respectively). This is likely a
reflection of rapid initial reaction followed by slower reaction of
the incorporated but unreacted (dangling) amine functionalities,
with associated network rearrangement, after the onset of the
chemical gelation. The even longer timescales of chemical gelation
and completion of crosslinking at 18.degree. C. was simply a
reflection of lower rates of reaction and diffusion at the lower
temperature.
After equilibrium swelling analysis, hydrogels were allowed to
degrade at 37.degree. C. in pH 7.4 PBS. FIG. 33 shows the rapid
mass loss of the hydrogels associated with the degradation of the
PAMAM crosslinkers beginning at seven and terminating at ten weeks.
While this is likely to be an appropriate degradation timescale for
many tissue engineering applications, it is worth pointing out once
more that one of the major advantages of this novel class of
polyamidoamine-based injectable hydrogels is the documented ability
to easily tune the rate of polymer degradation as needed through
alteration of the PAMAM backbone. Additionally, the hydrogels were
evaluated under base-catalyzed accelerated degradation conditions,
whereby 1 mL of 0.1N sodium hydroxide solution was added to each
vial in place of PBS. Within 4 days, hydrolytic degradation of the
PAMAM crosslinkers was completed in all gels to yield clear liquid
solutions at 4.degree. C. (data not shown).
Example 12
Cell Culture Studies & Cytocompatibility of Hydrogel
Leachables
A rat fibroblast cell line (ATCC, CRL-1764) was cultured on T-75
flasks using Dulbecco's modified Eagle medium (DMEM; Gibco Life,
Grand Island, N.Y.) supplemented with 10% (v/v) fetal bovine serum
(FBS; Cambrex BioScience, Walkersville, Md.) and 1% (v/v)
antibiotics containing penicillin, streptomycin and amphotericin
(Gibco Life). Cells were cultured in a humidified incubator at
37.degree. C. and 5% CO.sub.2. Cells of passage number 3 were used
in this study.
The cytocompatibility of the chemically and thermally gelled
hydrogels were evaluated by a leachables extraction test, according
to established protocols. (Klouda 2009 and Timmer 2003). Hydrogel
discs were formed by injecting 90 .mu.L aliquots of a 4.degree. C.
solution with 7 and 10 wt % PAMAM and TGM, respectively, in cell
culture media (DMEM, supplemented with antibiotics) without the
addition of serum into Teflon molds (6 mm in diameter and 3 mm in
height) held at 37.degree. C. and given 24 h to equilibrate.
Hydrogels were then immersed in an excess of cell culture media
without serum at a surface area to fluid volume ratio of 3
cm.sup.2/mL and incubated at 37.degree. C. for 24 h. (Timmer 2003).
The resulting hydrogel leachables solution was collected, sterile
filtered, and prepared in 1, 10, and 100.times. dilutions. Cultured
cells were harvested at 80-90% confluency with a Trypsin/EDTA
solution (2 mL/flask), resuspended at a density of 100,000
cells/mL, and seeded into 96-well tissue culture plates (100 .mu.L
cell suspension/well) for a seeding density of 10,000 cells/well.
The plates were then incubated for 24-48 h before testing to
achieve 80-90% confluency within the well. The three dilutions of
extract media were added to the cultured fibroblast cells in the
96-well plates (100 .mu.L/well), replacing the culture media (n=6).
In addition, cells fed with identical media without the extracted
leachables (DMEM supplemented with antibiotics and without serum)
served as a positive (live) control, and cells exposed to 70%
ethanol for 10 min served as the negative (dead) control (n=6). The
plates were then incubated at 37.degree. C., 95% relative humidity,
and 5% CO.sub.2 for either 2 or 24 h.
Following incubation, media were removed, the cells were rinsed
three times with pH 7.4 PBS, calcein AM and ethidium homodimer-1 in
2 .mu.M and 4 .mu.M concentrations in PBS, respectively (Live/Dead
viability/cytotoxicity kit, Molecular Probes, Eugene, Oreg.), were
added, and the cells were incubated in the dark at room temperature
for 30 min. Cell viability was then quantified using a fluorescence
plate reader (Biotek Instrument FLx800, Winooski, Vt.) equipped
with filter sets of 485/528 nm (excitation/emission) for calcein AM
(live cells) and 528/620 nm (excitation/emission) for EthD-1 (dead
cells). The fluorescence of the cell populations was recorded and
the fractions of live and dead cells were calculated relative to
the controls. The data are expressed as mean.+-.standard deviation,
and statistically significant differences were determined by
Tukey's post hoc test.
Finally, initial cytocompatibility of the hydrogels was evaluated
through exposure of rat fibroblasts to the soluble hydrogel
leachables with fluorescent Live/Dead analysis after 2 and 24 h.
FIG. 34 shows the resulting fraction of live cells treated with 1,
10 and 100.times. dilutions of the hydrogel leachables in media
relative to the untreated live control. The hydrogels were shown to
be fully cytocompatible at all conditions tested, which was largely
expected from the wealth of literature showing the general
cytocompatibility of PNiPAAm-based and polyamidoamine-based
macromers, as discussed in the introduction.
Example 13
Biocompatibility Evaluation of Poly(N-isopropylacrylamide)-based
Hydrogels for Craniofacial Bone Regeneration
The objectives of this study were (i) to fabricate non-shrinking,
biodegradable hydrogels by copolymerizing the PNiPAAm-based
macromers with pendant lactone rings to enable hydrolysis-dependent
degradation via LCST modulation and crosslinking with
polyamidoamine (PAMAM) crosslinkers and (ii) to evaluate the in
vitro cytocompatibility of the leachable and degradation byproducts
and the in vivo biocompatibility of the injectable system in an
orthotopic defect.
Thermogelling macromers (TGMs) were prepared with PNiPAAm, glycidyl
methacrylate, acrylic acid and the hydrolyzable ring,
dimethyl-.gamma.-butyrolactone acrylate (DBA), via free radical
polymerization by adapting the protocol as previously described
(Ekenseair A K. Biomacromolecules. 2012; 13 (6):1908-1915). Low
molecular weight PAMAM crosslinkers were created using a simple
polymerization following established protocols (Ekenseair A K.
Biomacromolecules. 2012; 13 (6):1908-1915). The cytocompatibility
of TGMs and crosslinked hydrogels was assessed with a fibroblast
cell line using leachable assays following previous studies (Klouda
L. Biomaterials. 2009; 30:4558-4566). Cell viability was quantified
using Live/Dead reagents and fluorescence plate reader and
normalized to controls. In vivo evaluation of two hydrogel
formulations (n=7) was performed in an 8 mm rat calvarial critical
size defect following established protocols (Spicer P. Nature
Protocols. 2012; 7:1918-1929). After harvest at 4 and 12 weeks,
samples were analyzed with microcomputed tomography (microCT),
histology and histomorphometry for biocompatibility, syneresis and
mineralization.
Rapid gelling, non-shrinking hydrogels were fabricated from the
mixing of TGMs with the PAMAM crosslinkers, resulting in highly
swollen gels. Extensive cytocompatibility testing of the TGM and
hydrogel demonstrated that hydrogel system presented little
cytotoxicity, and there were no significant effects of different
hydrogel parameters on cell viability except at the highest polymer
densities (FIG. 35). Additionally, the hydrogels did not impede
neotissue formation within the defect (FIG. 36).
The results indicate that the presence of dual thermal and chemical
crosslinking mechanisms can reduce hydrogel syneresis, which is
beneficial for the incorporation and proliferation of cells.
Furthermore, the hydrogel leachable products demonstrate in vitro
cytocompatibility and the preliminary data suggest these hydrogels
are biocompatible and potentially mineralize in vivo. In
combination with MSCs, this in situ forming hydrogel system may
provide a novel solution for localized and minimally invasive cell
delivery for craniofacial bone regeneration.
Example 14
Evaluation of Implanted Poly(N-isopropylacrylamide)-based Hydrogels
in Rat Cranial Defect
The objective of this study was to evaluate the mineralization
capacity and biocompatibility of acellular injectable, dual-gelling
hydrogels based on PNiPAAm for the healing of an 8 mm critical size
rat cranial defect.
Materials and Methods:
Materials
N-isopropylacrylamide (NiPAAm), dimethyl-.gamma.-butyrolactone
acrylate (DBA), glycidyl methacrylate (GMA), acrylic acid (AA),
2,2'-azobis(2-methylpropionitrile) (azobisisobutyronitrile, AIBN),
N,N'-methylenebisacrylamide (MBA), and piperazine (PiP) were
purchased from Sigma Aldrich (Sigma, St. Louis, Mo.) and used as
received. Anhydrous 1,4-dioxane, dimethylformamide, diethyl ether,
and acetone in analytical grade; water, acetonitrile, chloroform,
and methanol in HPLC-grade; and 1 N sodium hydroxide (NaOH) were
purchased from VWR (Radnor, Pa.) and used as received. Phosphate
buffered saline (PBS) solution was obtained from Gibco Life, Grand
Island, N.Y. (powder, pH 7.4). Ultrapure water was obtained from a
Millipore Super-Q water system (Millipore, Billerica, Mass.).
In Vivo Experimental Design
Three groups of acellular hydrogels were examined in an 8 mm
critical size rat cranial defect, as outlined in Table 3. The first
experimental group ("15 wt % TGM") consisted of the
P(NiPAAm-co-GMA) TGM at 15 wt % polymer. The other two experimental
groups ("15 wt % TGM/DBA") and ("20 wt % TGM/DBA") consisted of the
P(NiPAAm-co-GMA-co-DBA-co-AA) TGM with the DBA-containing
hydrolyzable lactone ring at 15 or 20 wt % polymer. Based on
previous work with rat cranial defects, the time periods chosen
were 4 and 12 weeks. At both timepoints, the samples were evaluated
with microCT and histology after harvest. All groups consisted of
n=6-7 rats.
TABLE-US-00003 TABLE 3 Study design for in vivo study Thermogelling
Macromer Polyamidoamine (TGM) (PAMAM) Crosslinker Group TGM Initial
Peak PAMAM Amine:epoxy TGM composition wt % LCST (.degree. C.)
M.sub.n(Da) mol ratio 1 P(NiPAAm.sub.92.6-co-GMA.sub.7.4) 15 30.9
.+-. 0.4 1440 1:1 P(NiPAAm.sub.84.9-co-GMA.sub.9.1- 2
co-DBA.sub.5.8-co-AA.sub.3.2) 15 22.4 .+-. 1.0 1440 1:1
P(NiPAAm.sub.84.9-co-GMA.sub.9.1- 3 co-DBA.sub.5.8-co-AA.sub.3.2)
20 22.4 .+-. 1.0 1440 1:1 *subscripts indicate mol % of
comonomer
TGM Synthesis and Characterization
Synthesis of P(NiPAAm-co-GMA) and P(NiPAAm-co-GMA-co-DBA-co-AA)
TGMs were performed according to those protocols reported in
Example 4 and as summarized herein below. In a typical reaction, 20
g of NiPAAm, GMA, DBA and AA was dissolved in 200 mL of either
anhydrous dimethylformamide or 1,4-dioxane under nitrogen at
65.degree. C. for the TGM and TGM/DBA groups, respectively. AIBN
pre-dissolved in the solvent was added at 0.7% of total mol content
to thermally initiate free radical polymerization, and the reaction
mixture was stirred for 16 h. After solvent removal by rotary
evaporation, the material was re-dissolved in either pure acetone
or a 95:5 (v/v) mixture of acetone:methanol for the TGM/DBA and TGM
groups, respectively and purified twice via dropwise precipitation
in at least 10.times. excess diethyl ether. The recovered polymer
was air-dried overnight and transferred to a vacuum oven for
several days prior to elemental analysis. The chemical composition
of the TGMs was determined by proton nuclear magnetic resonance
spectroscopy (1H NMR, Bruker, Switzerland). The polymer was
dissolved in D2O at a concentration of 20 mg/mL that contained 0.75
wt % 3-(trimethylsilyl)propionic-2,2,3,3-d4 acid, sodium salt as an
internal shift reference (Sigma-Aldrich, St. Louis, Mo.) and the
data were analyzed using the MestRe-C NMR software package
(Mestrelab Research S.L., Spain). Acid titration was performed in
conjunction with 1H NMR to determine the AA content of the TGMs
before hydrolysis. Aqueous gel permeation chromatography (GPC)
using a Waters Alliance HPLC system (Milford, Mass.) and
differential refractometer (Waters, model 410) equipped with a
series of analytical columns (Waters Styragel guard column 20 mm,
4.6.times.30 mm; Waters Ultrahydrogel column 1000, 7.8.times.300
mm) was used to determine the molecular weight distributions of the
synthesized TGM/DBA polymers. The TGM/DBA polymer was first
hydrolyzed in accelerated conditions to remove its thermogelling
properties. The weight average molecular weight (Mw), number
average molecular weight (Mn), and polydispersity index (PDI=Mw/Mn)
of the hydrolyzed polymer were determined by comparison to
commercially available narrowly dispersed molecular weight
poly(ethylene glycol) (PEG) standards (Waters, Mississauga, ON).
The LCSTs of the TGMs were determined by differential scanning
calorimetry (DSC). 14 .mu.L, of each polymer solution was pipetted
into an aluminum volatile sample pan (TA Instruments, Newcastle,
Del.) and capped/crimped. Thermograms were recorded on a TA
Instruments DSC 2920 equipped with a refrigerated cooling system
against an empty sealed pan as reference. The oven was equilibrated
at -5.degree. C. for 10 min and then heated to 80.degree. C. at a
heating rate of 5.degree. C./min. The LCST was determined both as
the onset and peak temperature of the endothermic peak in the
thermogram using the Universal Analysis 2000 software provided with
the DSC system. For this study, P(NiPAAm92.6-co-GMA7.4) with a
Mn=.about.9.2 kDa and PDI of 3 and
P(NiPAAm84.9-co-GMA9.1-co-DBA5.8-co-AA3.2) with a Mn=.about.56.1
kDa and PDI of 2 were used.
PAMAM Synthesis and Characterization
PAMAM was synthesized by the polyaddition of piperazine (PiP) and
methylene bisacrylamide (MBA) at a stoichiometric ratio of
[MBA]/[PiP]=0.75 as described previously herein (Example 5).
Molecular weight distributions of the synthesized PAMAM
crosslinkers were analyzed using time-of-flight mass spectroscopy
with positive-mode electrospray ionization on a Bruker microTOF ESI
spectrometer (Bruker Daltonics, Billerica, Mass.) equipped with a
1200 series HPLC (Agilent Technologies, Santa Clara, Calif.) to
deliver the mobile phase (50:50 HPLC-grade water and methanol).
After data acquisition, all peaks (including degradation and
secondary reaction products) were identified using microTOF Control
software (Bruker). The peaks were corrected for charge state
(generally with H.sup.+ or Na.sup.+ and rarely K.sup.+ ions), and
quantified for calculation of M.sub.n, M.sub.w, and PDI. PAMAM with
M.sub.n=1440 Da and PDI=1.38 was used for this study.
Hydrogel Fabrication
Hydrogels (8 mm in diameter, 2 mm in height) were fabricated by
combining the TGM and PAMAM crosslinker. The TGM and PAMAM
crosslinker were first sterilized via UV-irradiation for 2 h.
Individual solutions of TGM and PAMAM crosslinker were then
prepared at twice the desired concentrations (Table 3) in sterile
PBS pH 7.4 and placed on a shaker table at 4.degree. C. until
dissolved. Under sterile conditions, the PAMAM solution was
pipetted into the TGM solution using cold pipet tips and the
resulting solution was manually mixed in the glass vial. 110 .mu.L
injections were transferred to 8 mm diameter.times.2 mm height
cylindrical Teflon molds at 37.degree. C. either immediately for
TGM/DBA groups or after a 20 minute delay for the TGM group,
covered with a glass slide, and subsequently allowed to gel for 24
h prior to implantation.
Animal Surgeries and Euthanasia
This work was done in accordance with protocols approved by the
Rice University Institutional Care and Use Committee. After
hydrogel fabrication with PBS pH 7.4 under sterile conditions in 8
mm.times.2 mm Teflon molds, the hydrogels were implanted within an
8 mm rat cranial defect. All scaffold groups were statistically
randomized to minimize operative- or animal-related error. 11-12
week Fischer 344 rats weighing 176-200 g (Harlan, Indianapolis,
Ind.) were placed under 4% isoflurane and maintained at 2%
isoflurane/O.sub.2 gas mixture for the duration of the operation.
The incision site and surrounding area were shaved, sterilized with
povidine-iodine swabs, and subcutaneously injected with 500 .mu.L
of 1% lidocaine for local anesthesia. A linear incision was
performed from the nasal bone to the mid-sagittal crest, and the
skin and periosteum were exposed from the underlying bone. An 8 mm
craniotomy was performed with a dental surgical drilling unit with
a trephine burr and saline irrigation, and the calvarial disk was
carefully removed to prevent dural tearing. After cleaning and
removal of bone fragments, the scaffold was implanted and the
periosteum and skin were each separately closed with interrupted
braided Vicryl stitches. The rats were given an intraperitoneal
injection of saline (1 mL/100 g/h of anesthesia) to counteract
blood loss and aid recovery. Additionally, intraperitoneal
injections of buprenorphine (0.05 mg/kg) were given at 12, 24, and
36 h as post-operative analgesia. After surgery, the rats were
placed under pure O.sub.2 until awakened from anesthesia and
individually housed in soft-bedding cages. Animals were given free
access to food and water and monitored for complications. The
implants and surrounding tissue were harvested 4 or 12 weeks
post-surgery. The rats were euthanized by CO.sub.2 inhalation after
anesthesia under 4-5% isoflurane, followed by a bilateral
thoracotomy.
Implant Retrieval
The implants were harvested by making an incision between the
medial canthi of the eyes down to the bone using a 701 burr
attached to a Stryker Total Performance System straight handpiece
at 40,000 rpm with water irrigation. Similar cuts were made along
the left and right temporal bone and posterior aspect of the
cranial vault, resulting in a rectangular section of the cranium
containing the defect site and implant. The implants were fixed in
10% formalin (Formalde-Fresh, Fisher) for 3 days at 37.degree. C.
and transferred to 70% (v/v) ethanol for microCT and histological
analysis.
Microcomputed Tomography
MicroCT analysis with a Skyscan 1172 High-Resolution Micro-CT
(Aartselaar, Belgium) with 10 .mu.m resolution, 0.5 mm aluminum
filter, and voltage of 100 kV and current of 100 .mu.A was used to
examine the morphology and mineralization of the implants and bony
union across the defect. Volumetric reconstruction and analysis was
conducted using Nrecon and CT-analyser software provided by
Skyscan. The percent of bone formation, including hydrogel
mineralization, within the defect was determined by centering a
cylindrical volume of interest (VOI) of 8 mm in diameter and 2 mm
in height at the bottom of the defect. Data are reported as the %
binarized object volume measured within this VOI within
thresholding gray values (70-255) with the CT-analysis software as
described in Henslee, A. M., et al., Acta Biomater 7 (10): p.
3627-37. The extent of bony bridging and union within the defect
were scored according the grading scale in Table 4. These scores
were determined from maximum intensity projections of the samples
generated from the microCT datasets. 3D models of each sample were
also generated to visualize the distribution of mineral deposits
within the hydrogel implants.
Histological Processing
After microCT scanning, the samples were sent to the MD Anderson
Bone Histomorphometry Core laboratory for dehydration and embedding
in poly(methyl methacrylate). 5 .mu.m coronal cross-sections were
taken from the center of the defect of each sample and staining was
performed with von Kossa, hematoxylin and eosin, and Goldner's
trichrome.
Histological Scoring
The three histological sections were evaluated via light microscopy
(Eclipse E600, Nikon, Melville, N.Y. with attached 3CCD Color Video
Camera DXC-950P, Sony, Park Ridge, N.J.) and scored using
histological scoring analysis. The histological evaluation was
performed along random implant-tissue interfaces within central
coronal cross-sections to assess: (1) overall tissue response
following the rubric outlined in Table 6 in the Results below, (2)
mineralization within the hydrogel following the scoring guide
outlined in Table 5, and (3) extent of bony bridging across the
defect following the rubric outline in Table 5. The evaluations
were performed blindly with randomized samples by three reviewers.
A description of the scoring system for mineralization and bony
bridging is shown in Table 5.
TABLE-US-00004 TABLE 5 Mineralization of Scaffold and Bony Bridging
Scoring Guide Description Score Mineralization of Scaffold Mineral
deposition in 75-100% of the scaffold 4 Mineral deposition in
50-75% of the scaffold 3 Mineral deposition in 25-50% of the
scaffold 2 Mineral deposition in up to 25% of the scaffold 1 No
mineral deposition observed 0 Bony Bridging Across Full Length of
Scaffold >75% bridging across defect 4 50-75% bridging across
defect 3 25-50% bridging across defect 2 <25% bridging across
defect 1 No bony bridging observed 0
Acellular Mineralization
8.times.2 mm acellular hydrogels (n=4 per group) were fabricated as
described above. The hydrogels were incubated in complete
osteogenic medium without fetal bovine serum (FBS), complete
osteogenic medium with 10% FBS (Cambrex Bioscience), PBS pH 7.4,
and 1.times. simulated body fluid for 0, 1, 7, 14, and 28 days at
37.degree. C. in 12-well plates. The simulated body fluid was
prepared according to Kokubo, T. and H. Takadama, Biomaterials,
2006. 27 (15): p. 2907-15. Each solution was changed every 2-3
days. At timepoints, the hydrogels were soaked in ultrapure water
for 30 min, cut in half, blotted and weighed. Sample halves (n=4
halves) for the calcium assay were place in 500 .mu.L of
ddH.sub.2O, homogenized through three freeze/thaw/sonication
cycles, and digested overnight in equal parts of 1N acetic acid for
a final concentration of 0.5 N acetic acid. The assay was performed
with a commercially available kit (Sekisui Diagnostics, Tokyo,
Japan) according to the manufacturer's instructions. The remaining
sample halves were fixed in 10% formalin, dehydrated in 70%
ethanol, placed in Histoprep frozen tissue embedding media (Fischer
Scientific, Waltham, Mass.), and frozen at -20.degree. C.
Statistics
The microCT and histological scoring data were analyzed via one-way
analysis of variance followed by the Kruskal-Wallis test
(p<0.05) for n=6-7 samples. The microCT bone volume for n=6-7
samples and in vitro mineralization data for n=4 were presented as
means.+-.standard deviation, unless otherwise stated. The acellular
mineralization data were analyzed by Tukey's post-hoc test using
JMP v11 statistical software.
Results:
MicroCT analysis
Three different groups of acellular, dual-gelling scaffolds were
implanted: one group without the DBA monomer containing a
hydrolyzable lactone ring for LCST modulation (15 wt % TGM), and
two groups with the DBA monomer at different polymer wt %
concentrations (15 and 20 wt % TGM/DBA). MicroCT was used for
nondestructive, quantitative analysis of bony bridging, union, and
bone volume, as well as 3D visualization of the extent and spatial
distribution of mineralization. For bony bridging of the defects,
maximum intensity projections (MIPs) were generated from each of
the microCT datasets and scored by three blinded reviewers using
the 0-4 grading scale detailed in Table 4. FIG. 43 provides
respective examples of actual MIPs for each score (A=4; B=3; C=2;
D=1). FIG. 37 shows the results from the scoring at 4 and 12 weeks.
At the 4 week timepoint, the average union microCT score for 15 wt
% TGM, 15 wt % TGM/DBA, and 20 wt % TGM/DBA was .about.2,
indicating that bony bridging is only observed at the defect
borders. At the 12 week timepoint, the 15 wt % TGM and 20 wt %
TGM/DBA demonstrated an increase in bony bridging, leading to a
statistically significant increase in the average union microCT
score compared to the 4 week timepoint.
TABLE-US-00005 TABLE 4 Scoring guide for bony bridging and union in
microCT datasets Description Score Bony bridging over entire defect
span at longest point 4 Bony bridging over partial length of defect
3 Bony bridging only at defect borders 2 Few bony spicules
dispersed throughout defect 1 No bone formation within defect 0
FIG. 38 shows the results from the quantification of bone volume
using microCT. The data are reported as the % binarized object
volume within a given VOI above a critical threshold grayscale
value of 70, and include mineralization within the implant and bone
formation across the defect. Both the 15 wt % TGM and 20 wt %
TGM/DBA groups demonstrated a significant increase in the bone
volume from 4 to 12 weeks. Additionally, the bone volume of the 15
wt % TGM group was significantly higher at 12 weeks compared to the
two other groups.
To examine the spatial distribution and nature of mineralization in
detail, 3D models of each sample were generated via microCT using
the critical threshold value. Two representative samples in the 15
wt % TGM and 20 wt % TGM groups demonstrating the most
mineralization at 12 weeks are shown in FIG. 39. Bone regeneration
across the defect at 12 weeks primarily extended underneath the
hydrogel for both groups. Mineralization additionally occurred
throughout the hydrogel for the 15 wt % TGM group. This is in
contrast to other samples, which showed bone formation at the
defect borders, or, as demonstrated by the representative 4 week
sample in FIG. 39, mineralization as nodules above the
hydrogel.
Descriptive Light Microscopy
The samples were grossly examined after histological staining with
hematoxylin & eosin (H&E), von Kossa, and Goldner's
trichrome, and representative images from each group and time point
are shown in FIG. 40. For all sections, the implants were visible,
although shrinkage and delamination artifacts generally associated
with hydrogel histology were observed in many samples. For both the
4 and 12 week samples, a thin fibrous capsule was observed around
the hydrogel, particularly around the periosteal border. At 4
weeks, the tissue at the hydrogel-defect interface was generally
loose and fibrous in nature and minimal bone formation was
observed. The inflammatory response, characterized by the presence
of neutrophils and macrophages, was also minimal and occurred
primarily at the periosteal border. At 12 weeks, the inflammatory
response was generally mitigated and the tissue at the
hydrogel-defect interface was more organized, as shown by the
high-magnification subsets in FIG. 40. Although all the hydrogels
did not demonstrate measurable scaffold fragmentation, instances of
cell-mediated bone formation across the defect were more pronounced
in the 15 wt % TGM group compared to the other groups. With H&E
and Goldner's trichrome staining, cell-mediated bone formation was
generally observed on the dural side of the defect, although in
certain 15 wt % TGM samples, significant mineralization was also
seen throughout the center of the hydrogel with von Kossa staining.
Additionally, promising direct bone-implant contact was observed in
several of the 15 wt % TGM samples at 12 weeks, a representative
example of which is shown in FIG. 40 (panel F) and the
high-magnification subsets. With the coronal cross sectional
slices, no significant amounts of hydrogel mineralization within
the implant or cell-mediated bone formation across the dural side
of the defect were observed.
Quantitative Histological Analysis
Samples were evaluated for bony bridging across the defect and
mineralization of the scaffold as shown in FIG. 42A-B according to
the inlaid scoring rubrics. FIG. 42A shows the scoring results for
mineralization within the scaffold. The average score for the 15 wt
% TGM group at 4 and 12 weeks was higher than that of the 15 wt %
TGM/DBA and 20 wt % TGM/DBA groups, but not significantly. FIG. 42B
shows the scoring results for bony bridging across the central
coronal cross-section of the scaffold along the dural side of the
implant. The 15 wt % TGM group showed significantly increased
scores in bony bridging from 4 to 12 weeks, but the scores were
only significant from the 15 wt % TGM/DBA group at 12 weeks. The
overall tissue response was also scored (Table 6), with only the 20
wt % TGM group showing a significant difference between the 4 and
12 week timepoints.
TABLE-US-00006 TABLE 6 Histological scoring guide and scores for
overall tissue response Overall Tissue Response Description Score
Direct bone to implant contact without soft interlayer 4 Remodeling
lacuna with osteoblasts and/or osteoclasts at surface 3 Majority of
implant is surrounded by fibrous tissue capsule 2 Unorganized
fibrous tissue (majority of tissue is not arranged as 1 capsule)
Inflammation marked by an abundance of inflammatory cells and 0
poorly organized tissue Timepoint Average Histological Group
(weeks) Score 4 0.14 .+-. 0.4 15 wt % TGM 12 0.83 .+-. 1.2 4 0.29
.+-. 0.5 15 wt % TGM/DBA 12 0.86 .+-. 0.7 4 0.00 .+-. 0.0 20 wt %
TGM/DBA 12 0.83 .+-. 0.8* *significant difference between the 4 and
12 week timepoint (p < 0.05)
In Vitro Acellular Mineralization
To investigate the mechanism behind hydrogel mineralization, an
acellular in vitro study was performed by culturing the hydrogels
in four solutions: PBS pH 7.4 (PBS), SBF 1.times. (SBF), complete
osteogenic media without serum (NS), and complete osteogenic media
with serum (S). FIG. 41 shows the calcium content of the hydrogels
over the 28 day culture period. All of the groups demonstrated
significantly increased calcium content over time after culture in
complete osteogenic media with serum. This corresponded with
increased von Kossa histological staining of hydrogel
cross-sections over time for all groups, which detects the presence
of phosphate groups (data not shown). However, the amount of
calcium in the 15 wt % TGM group was significantly larger than that
of the 15 wt % and 20 wt % TGM/DBA groups at the 7, 14 and 28 day
timepoints. The 15 wt % TGM group also displayed significantly
increased calcium content in the SBF 1.times. and complete
osteogenic media without serum (NS) conditions; however, these
values were not statistically different from the calcium content of
complete osteogenic media with serum (S) conditions.
Conclusion:
This Example evaluated the mineralization capacity, bone formation
capability, and tissue response of acellular PNiPAAm-based
hydrogels in an 8 mm critical size rat cranial defect. Although
minimal bone formation and hydrogel mineralization was observed at
4 weeks for all groups, significantly higher bone formation across
the defect and mineralization within the implant was observed for
the 15 wt % TGM and 20 wt % TGM/DBA groups after 12 weeks. The
results, coupled with an in vitro acellular mineralization study,
suggest that these hydrogels undergo matrix
hydrophobicity-dependent mineralization, which is accelerated by
the adsorption of calcium-binding and nucleating proteins on
hydrophobic hydrogel surfaces. These mineralizable and
biocompatible injectable hydrogels possess great potential as
acellular strategies or stem cell carriers for craniofacial tissue
engineering applications.
Therefore, the present disclosure is well adapted to attain the
ends and advantages mentioned as well as those that are inherent
therein. The particular embodiments disclosed above are
illustrative only, as the present disclosure may be modified and
practiced in different but equivalent manners apparent to those
skilled in the art having the benefit of the teachings herein.
Furthermore, no limitations are intended to the details of
construction or design herein shown, other than as described in the
claims below. It is therefore evident that the particular
illustrative embodiments disclosed above may be altered or modified
and all such variations are considered within the scope and spirit
of the present disclosure. While compositions and methods are
described in terms of "comprising," "containing," or "including"
various components or steps, the compositions and methods can also
"consist essentially of" or "consist of" the various components and
steps. All numbers and ranges disclosed above may vary by some
amount. Whenever a numerical range with a lower limit and an upper
limit is disclosed, any number and any included range falling
within the range is specifically disclosed. In particular, every
range of values (of the form, "from about a to about b," or,
equivalently, "from approximately a to b," or, equivalently, "from
approximately a-b") disclosed herein is to be understood to set
forth every number and range encompassed within the broader range
of values. Also, the terms in the claims have their plain, ordinary
meaning unless otherwise explicitly and clearly defined by the
patentee. Moreover, the indefinite articles "a" or "an," as used in
the claims, are defined herein to mean one or more than one of the
element that it introduces. If there is any conflict in the usages
of a word or term in this specification and one or more patent or
other documents that may be incorporated herein by reference, the
definitions that are consistent with this specification should be
adopted.
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